The field of the present invention is computed tomography and, particularly, computer tomography (CT) scanners used to produce medical images from X-ray attenuation measurements.
As shown in FIG. 1, a CT scanner used to produce images of the human anatomy includes a patient table 10 which can be positioned within the aperture 11 of a gantry 12. A source of highly collimated X-rays 13 is mounted within the gantry 12 to one side of its aperture 11, and one or more detectors 14 are mounted to the other side of the aperture. The X-ray source 13 and detectors 14 are revolved about the aperture 11 during a scan of the patient to obtain X-ray attenuation measurements from many different angles about the patient.
A complete scan of the patient is comprised of a set of X-ray attenuation measurements which are made at discrete angular orientations of the X-ray source 13 and detector 14. Each such set of measurements is referred to in the art as a "view" and the results of each such set of measurements is a transmission profile. As shown in FIG. 2, the X-ray source 13 produces a fan-shaped beam which passes through the patient and impinges on an array of detectors 14. Each detector 14 in this array produces a separate attenuation signal and the signals from all the detectors 14 are separately acquired to produce the transmission profile for the indicated angular orientation. The X-ray source 13 and detector array 14 are then revolved in direction 15 to a different angular orientation where the next transmission profile is acquired.
As the data is acquired for each transmission profile, the signals are sampled, filtered and stored in a computer memory. The signal from the detectors are oversampled to provide twice the number of transmission profiles as are required to reconstruct an image, for example. The attenuation measurement samples then are digitally low-pass filtered and the output of the filtering is sampled at a rate that produces the required number of transmission profiles from which to reconstruct an image. These steps are performed in real time as the data is being acquired.
The resultant transmission profiles then are used to reconstruct an image which reveals the anatomical structures in a slice taken through the patient. The prevailing method for reconstructing image is referred to in the art as the filtered back projection technique. The attenuation measurements are converted to integers called "CT numbers" or "Hounsfield units", which are used to control the brightness of a corresponding pixel on a CRT display.
Each X-ray detector 14 comprises a scintillator and a solid state photodiode. X-rays striking the scintillator produce light photons which are absorbed by the photodiode creating an electric current. The light is not emitted by the scintillators instantaneously, rather the emission follows a multi-exponential decay curve. In fact, the time dependence of the emitted light intensity can be modelled accurately as a sum of exponentials with different decay constants. The term "primary speed" refers to the most prompt of these exponential decays and is defined as the time in which the detector light output falls to 1/e of its initial value after being stimulated by an impulse of X-rays.
As the detector array is rapidly rotating about the patient, the exponential decay blurs together detector readings for successive views. This blurring, due to the response time lag of the detector, is referred to as "afterglow" and degrades the azimuthal component of the image resolution. The azimuthal direction 16 of the image area is perpendicular to a line 17 from the center of the imaging aperture 11. The amount of blurring increases the farther the object is spaced from the aperture center, since the speed at which the object is swept across the detectors 14 effectively increases with this spacing.
FIG. 3 plots attenuation values from a given detector for a series of views and graphically depicts the blurring. The solid line represents the output of a single detector 14 during several views for a square object being imaged. Ideally the detector data should have a pulse-like shape as represented by the dashed lines. However, the effect of the afterglow blurring rounds the edges of the waveform and extends the object signal into several adjacent views. When the views are used to reconstruct an image, the object will appear enlarged and will not have sharp, distinct edges.
An obvious solution to the resolution degradation is to slow the rotational speed of the x-ray source and detectors. However, this prolongs image acquisition and discomfort to the patient. Heretofore, a certain level of degradation has been tolerated, but as rotational scan periods become shorter, approaching one second for example, the image degradation reaches unsatisfactory levels.
Another solution is to cause the detector to emit light quicker, e.g. build a better detector. As is well known in the art, the quantity of light emitted normally decreases with the speed of its emission. The decrease in the quantity of light results in greater statistical fluctuation in the detector output, causing a noisier image. It should be noted that this alternative solution would introduce noise into the image data to the same extent as the deconvolution technique of the present invention, when electronic noise is negligible compared to quantum noise.